A normal ear transmits sounds as shown in FIG. 1 through the outer ear 101 to the tympanic membrane (eardrum) 102, which moves the bones of the middle ear 103 (malleus, incus, and stapes) that vibrate the oval window and round window openings of the cochlea 104. The cochlea 104 is a long narrow duct wound spirally about its axis for approximately two and a half turns. It includes an upper channel known as the scala vestibuli and a lower channel known as the scala tympani, which are connected by the cochlear duct. The cochlea 104 forms an upright spiraling cone with a center called the modiolar where the spiral ganglion cells of the acoustic nerve 113 reside. In response to received sounds transmitted by the middle ear 103, the fluid-filled cochlea 104 functions as a transducer to generate electric pulses which are transmitted to the cochlear nerve 113, and ultimately to the brain.
Hearing is impaired when there are problems in the ability to transduce external sounds into meaningful action potentials along the neural substrate of the cochlea 104. To improve impaired hearing, auditory prostheses have been developed. For example, when the impairment is related to operation of the middle ear 103, a conventional hearing aid may be used to provide acoustic-mechanical stimulation to the auditory system in the form of amplified sound. Or when the impairment is associated with the cochlea 104, a cochlear implant with an implanted stimulation electrode can electrically stimulate auditory nerve tissue with small currents delivered by multiple electrode contacts distributed along the electrode.
FIG. 1 also shows some components of a typical cochlear implant system which includes an external microphone that provides an audio signal input to an external signal processor 111 where various signal processing schemes can be implemented. The processed signal is then converted into a digital data format, such as a sequence of data frames, for transmission into the implant processor 108. Besides receiving the processed audio information, the implant processor 108 also performs additional signal processing such as error correction, pulse formation, etc., and produces a stimulation pattern (based on the extracted audio information) that is sent through an electrode lead 109 to an implanted electrode array 110. Typically, this electrode array 110 includes multiple electrode contacts 112 on its surface that provide selective stimulation of the cochlea 104.
Cochlear implant systems employ stimulation strategies that provide high-rate pulsatile stimuli to electrode contacts in multi-channel electrode arrays. One specific example is the “Continuous Interleaved Sampling (CIS)”—strategy, as described by Wilson et al., Better Speech Recognition With Cochlear Implants, Nature, vol. 352:236-238 (1991), which is incorporated herein by reference. For CIS, symmetrical biphasic current pulses are used, which are strictly non-overlapping in time. The rate per channel typically is higher than 800 pulses/sec. Other stimulation strategies may be based on parallel activation of electrode currents. These approaches have proven to be successful in giving high levels of speech recognition.
For an audio prosthesis such as a cochlear implant to work correctly, some patient-specific operating parameters need to be determined in a fit adjustment procedure where the type and number of operating parameters are device dependent and stimulation strategy dependent. Possible patient-specific operating parameters for a cochlear implant include:                THR1 (lower detection threshold of stimulation amplitude) for Electrode Contact 1        MCL1 (most comfortable loudness) for Electrode Contact 1        Phase Duration for Electrode Contact 1                    Amplitude for Electrode Contact 1            Pulse Rate for Electrode Contact 1                        THR2 for Electrode Contact 2        MCL2 for Electrode Contact 2        Phase Duration for Electrode Contact 2                    Amplitude for Electrode Contact 2            Pulse Rate for Electrode Contact 2                        . . .        Number of fine structure channels        Compression        Parameters of frequency, e.g. electrode contact mapping        Parameters describing the electrical field distribution, e.g. spatial spread        
One approach for an objective measurement of MCLs and THLs is based on the measurement of the eCAPs (Electrically Evoked Compound Action Potentials), as described by Gantz et al., Intraoperative Measures of Electrically Evoked Auditory Nerve Compound Action Potentials, American Journal of Otology 15 (2):137-144 (1994), which is incorporated herein by reference. In this approach, a recording electrode in the scala tympani of the inner ear is used. The overall response of the auditory nerve to an electrical stimulus is measured very close to the position of the nerve excitation by a given electrode contact. This neural response is caused by the super-position of single neural responses at the outside of the axon membranes. The amplitude of the EAP at the measurement position is between 10 μV and 1800 μV. Other objective measurement approaches are also known, such as electrically evoked stapedius reflex thresholds (eSRT).
One common method for fit adjustment is to behaviorally find the threshold (THR) and most comfortable loudness (MCL) value for each separate electrode contact. For this, the stimulation charge on a selected electrode channel is usually increased in steps from zero until the THR or MCL level is reached in a subjective procedure (e.g. method of adjustments) or an objective procedure (e.g., eCAP or eSRT). This increase can be either stimulation burst duration or stimulation burst amplitude or a combination thereof. Typically, for this procedure constant amplitude stimulation bursts with 10-1000 msec duration are utilized. See for example, Rätz, Fitting Guide for First Fitting with MAESTRO 2.0, MED-EL, Fürstenweg 77a, 6020 Innsbruck, 1.0 Edition, 2007. AW 5420 Rev. 1.0 (English_EU); incorporated herein by reference. Other alternatives/extensions are sometimes used with a reduced set of operating parameters; e.g. as suggested by Smoorenburg, Cochlear Implant Ear Marks, University Medical Centre Utrecht, 2006; U.S. Patent Application 20060235332; which are incorporated herein by reference. Typically each electrode channel is fitted separately without using the information from already fitted electrode channels. The stimulation charge on a given electrode contact typically is increased in steps from zero until the MCL (most comfortable loudness) is reached.
Several approaches currently are used to accelerate the fitting process. One approach is to use a flat map, i.e. use the same MCL or THR value on all electrode channels so that only one electrode channel needs to be fitted. But this approach allows no conclusion to be drawn about the perceptive status (high or less sensitive) of fitted electrode channels and consequently the resulting map can be in much too loud or too soft for some electrode channels. Another approach is to increase electrode stimulation charge during fitting on N adjacent electrode contacts simultaneously from zero onwards and thereby so to speak fit N adjacent electrode contacts simultaneously. These and similar approaches do save time, however, they have the disadvantage of not taking into account electrode-specific particularities, like, e.g., a certain electrode channels having a considerably different MCL value from another electrode channels. A third used approach for example is to not start from zero when fitting an electrode channel, but from a certain fixed value. This approach however has the disadvantage of the fixed starting values possibly being too high with respect on MCL on one electrode channel and possibly being much too low with respect to MCL on another electrode channel. In other words, the risk of over-stimulating the patient exists, while there is still potential of more time savings.